1. Field of the Invention
The present invention relates generally to radiation detectors in nuclear medicine and more specifically to the electronics used to detect signals from solid state photosensors used in some radiation detectors.
2. Description of the Background Art
Nuclear medicine imaging is sensitive to the radionuclide distribution within a patient after the in vivo administration of radiopharmaceuticals to acquire images that are used to assess the function and anatomy of organs, bones or tissues of the body. The imaging systems that detect the radionuclide distribution comprise radiation detectors and associated electronics. The imaging systems detect x-ray or gamma ray photons derived from the administered radionuclides. Single photon emission imaging and coincidence imaging are two forms of nuclear medicine imaging that are currently in common use. In single photon emission imaging, the radionuclide itself directly emits the radiation to be detected. For example, in Single Photon Emission Computed Tomography (SPECT), gamma-emitting radionuclides such as 99Tc, 123I, 67Ga and 111In may be part of the administered radiopharmaceutical. The imaging system may use a lead collimator to eliminate all photons but those photons perpendicular to the surface of the detector. The location and energy of emitted photons may then be accumulated until a satisfactory image is obtained. Coincidence imaging eliminates the need for such a collimator by relying on the detection of two photons at different detectors at nearly the same time. An example of coincidence imaging in current clinical use is Positron Emission Tomography (PET). In PET, β+-emitting radionuclides such as 11C, 13N, 15O, 18F, 68Ga, 82Rb are part of the administered radiopharmaceutical. The emitted positrons react with electrons within the patient's body, the annihilation creating two 511 keV photons emitted in opposite directions. The two photons are then detected within a certain time window, generally in the nanosecond range, of each other. Radiation detectors for nuclear medicine imaging may need to detect photons from 1 keV to several MeV in energy.
Many radiation detectors in current use in nuclear medicine imaging systems consist of a scintillation crystal, or scintillator, for converting x-ray or gamma ray photons into visible light photons, so called scintillation photons, and a device for converting the scintillation photons into electrical signals. The Anger camera pioneered this approach in the 1950s and is more fully described in U.S. Pat. No. 3,011,057. The Anger camera consists of a NaI crystal and an array of photomultiplier tubes (PMTs). In operation, gamma ray photons cause scintillation photons to be emitted from the NaI crystal. The scintillator photons then impinge the different PMTs. The PMTs convert the detected light photons into electrical signals, which are amplified and inputted to a location and energy computing circuit, which yields information about the location and energy of the scintillation event within the NaI crystal.
Regardless of the scintillator, the scintillation photons produced must be converted into an electrical signal to be analyzed. A PMT is a vacuum tube including a photocathode, and an electron multiplier sealed into an evacuated glass tube, and an input window which is optically coupled to the scintillation crystal. Scintillation photons (4 or 5) incident on the photocathode cause the photocathode to emit an electron. The electron is absorbed by a dynode which emits 5-6 electrons. A series of dynodes repeat this reaction until a final large cluster of electrons is fed through the anode as a pulse to the attached logic circuits to determine position.
PMTs are extremely sensitive to low levels of light. However, PMTs have many drawbacks. PMTs require a high voltage for operation, typically greater than 1000V. PMTs are vulnerable to drifting in performance, especially early in their life cycle. PMTs are susceptible to mechanical failure and may thus be less reliable. PMTs are susceptible to magnetic fields, such as from the MRI devices (and even from the earth's comparatively weak magnetic field). PMTs are physically bulky. The size of the PMTs determines and limits the intrinsic spatial resolution of a detector system. Furthermore, PMTs require lead shielding, thus increasing the weight of the overall camera. This increases costs, especially in the case the camera must be moved by motors for tomographic imaging.
In addressing the above problems, photodetectors composed of an array of solid state photodiodes have been used rather than PMTs. See, for example, U.S. Pat. No. 5,171,998, incorporated by reference herein in its entirety. Inorganic photodiodes, generally various forms or compounds of silicon, address some of the problems of the PMTs. The inorganic photodiodes are more stable over their life cycle, more robust mechanically, not susceptible to magnetic fields, and much smaller and lighter. However, inorganic photodiodes are expensive, difficult, and slow to fabricate. Their mechanical structure is rigid. Inorganic photodiodes are susceptible to radiation damage. Inorganic photodiodes generally have a poor spectral response to long wavelength scintillation photons from certain scintillation crystals, such as CsI. Finally, the low band gap of silicon based photodiodes yields thermally generated leakage current, which acts as noise in associated circuits which read the signal from the inorganic photodiodes. The silicon photodiodes must be cooled to lower such leakage current to acceptable levels.
The use of carbon-based photodiodes in lieu of inorganic photodiodes has been disclosed. Using an array of carbon-based photodiodes in a photodetector has many advantages. One such advantage is a substantially lower leakage current (or“dark” current) than found in current inorganic photodiodes. Such substantially lower leakage current is due to the higher band gap seen in carbon-based photodiodes. However, the capacitance of such photodiodes is substantially increased.
Front-end electronics optimized for silicon photodiode detectors are designed to minimize noise and thus maximize energy resolution. Such front-end electronics thus cannot take advantage of the lower noise of the carbon-based photodiode and cannot realize the increase in energy yield due to the reduced noise from the leakage current. Front-end electronics that take advantage of the different electrical characteristics of low leakage current photodiodes, such as carbon-based photodiodes, have been proposed.
In particular, front-end electronics that use a number of shaper circuits with relatively long shaping or integration times have been proposed for use with low leakage current photodiodes.
FIG. 1 shows an example of such electronics. A radiation detector 2 includes a radiation detector module 4 and associated electronics 6. The radiation detector module 4 has a scintillator 8 and an array of photodiodes 10 (composed of individual photodiodes 12) optically coupled to the scintillator 8. The array of photodiodes 10 is electrically coupled to associated electronics 6. The associated electronics 6 include the front-end electronics 14 and the read-out electronics 16. The read-out electronics 16 is in communication with a processing apparatus such as computer 28. The front-end electronics 14 typically includes a preamplifier 20 for boosting the low output current of photodiode array 10. A shaper circuit 22 is coupled to the preamplifier circuit 20. The read-out electronics 16 includes a sample/hold circuit 24 for holding the shaped signal output from shaper circuit 22. Each individual photodiode 12 of photodiode array 10 represents a pixel of the image, and thus requires its own set of preamplifier, shaper, and sample/hold circuits that are the same as shown in FIG. 1. The sample/hold circuit 24 is triggered to send the held signal to multiplexer 26, which sorts signals from the multiple pixels of the array of photodiodes 10. The signal from the multiplexer 26 is then communicated to computer 28 or other logic circuits for computation of event location and energy as well as other calculations.
When an x-ray or gamma ray is absorbed within the scintillator 8, scintillation photons in the visible spectrum are generated. The scintillation photons which impinge a pixel 12 of the photodiode array 10 generate a current. The current is amplified by the preamplifier 20, and then integrated over a shaping time (ts) (also known as the “shaper peaking time” or “integration time”) by the shaper circuit 22. The resulting signal is a shaped pulse with peak amplitude related to the amount of charge accumulated over the integration time ts.
The duration of the current from the photodiode 12 depends on the detector decay time, which is an intrinsic detector property. This time is equivalent to the collection time (tc) of the current. However, making the integration time ts equal to the collection time tc is not sufficient to maximize the amount of signal collected from the current and converted into the pulse output of the shaper 22, due to the phenomenon of ballistic deficit. FIG. 2 shows the ballistic deficit for a CsI(TI)/photodiode detector. In this model, the shaper circuit is assumed to be of the first order. The detector decay time is assumed to be 1 μs. A shaping time ts equal to the collection time tc will yield a peak approximately equal to only one third of the signal received. In general, the shaping time should be at least five times larger than the collection time (ts≧5 tc) in order for the peak of the resulting shaped pulse to be in proportion to the energy of the absorbed photon collected over the collection time.
Thus, when lower leakage current (and higher capacitance) photosensors such as carbon-based photodiodes are used as the radiation detectors, several changes occur in the relevant electrical characteristics of the detector 4, which require changes in the front-end electronics 14 in order to minimize noise.
The thermal noise, which arises from the leakage (or dark) current, can be modeled as a parallel noise source. The “shot” noise, which is dependent on the detector capacitance, can be modeled as a series noise source. The increase in detector capacitance creates more series noise in the detector 4, but the decrease in dark current decreases the amount of parallel noise in detector 4. An analysis of the energy resolution versus shaping time as shown in FIG. 3 demonstrates that increasing the shaping time will yield an improved energy resolution.
According to the electronics of FIG. 1, either a trigger circuit is used to simultaneously trigger all the shaper circuits connected to the array of photodiodes, or the shaping circuits may be clocked at a high sample time ts such as 100 μs. This operation is based on the theory that the longer shaping time reduces the chance that a current pulse will not be fully integrated.
However, use of this method requires that the signals held in the sample/hold circuits 24 also be read out simultaneously. In other words, events are read out as they occur in time, one at a time, for the entire circuit. Such operation places a limitation on the maximum count rate obtainable by the system, as new event detection must be disabled during readout.